Sensor calibration and blood volume determination

ABSTRACT

A method and device for the continuous real-time monitoring of relative blood volume change, based on registration of blood hemoglobin concentration, during long periods of time, such as dialysis session. A method and device for cardiac output measurement during dialysis, surgeries, intensive care procedure is provided. Calibration of a sensor in response to a change in a measureable blood property is determined from a coefficient corresponding to a change in the measured blood property.

The present application claims priority as a continuation in part ofU.S. Ser. No. 09/528,880 filed Mar. 20, 2000, now U.S. Pat. No.6,493,567, which is a divisional of U.S. Ser. No. 08/950,244 filed Oct.14, 1997 now U.S. Pat. No. 6,041,246, each of which is hereby expresslyincorporated by reference.

FIELD OF THE INVENTION

The present invention relates to the photometric analysis of bloodproperties to monitor changes in blood volume, blood proteinsconcentration, cardiac output and other hemodynamic parameters. Moreparticularly, the invention includes a method and apparatus foremploying a single light emitter and a single photodetector, the singlephotodetector is oriented with respect to the emitter to minimize ascattering effect of the light from the electrolyte composition of theblood, blood flow rate, blood hematocrit level and other factors.

BACKGROUND OF THE INVENTION

The optical density of blood corresponds to a number of factors and themeasurement of the optical density of the blood has been used todetermine certain blood parameters.

For example, during a hemodialysis session when fluid is being removedfrom the blood stream by a dialyzing machine, the concentration ofhemoglobin, naturally occurring in the red blood cells, may decrease or,generally increase, with respect to the equilibrium processes of thefluid removal and mobilization from the body tissues. The change in theconcentration of hemoglobin results in a corresponding change in theoptical density of the blood and may be registered.

Monitoring changes in blood volume, as well as blood hemoglobin, canhelp to prevent such complications as hypotension due to continuoushypovolemia. Cardiac output of a patient, as well as other hemodynamicparameters of cardiovascular system, also provide useful informationabout the patient condition during hemodialysis, surgery, and inintensive care units. Using a dilution technique it is possible tocalculate several characteristics of cardiovascular system, particularlycardiac output.

There are several research works and patents in the field of photometricblood monitoring. As shown in U.S. Pat. No. 3,830,569 to Meric; U.S.Pat. No. 4,243,883 to Schwartzmann; and U.S. Pat. No. 4,303,336 toCullis, a device with suitable light source and photodetector may beused. Multiple detectors may be used, as shown by Meric.

As illustrated by U.S. Pat. No. 4,745,279 to Karkar et al.; U.S. Pat.No. 4,776,340 to Moran et al.; U.S. Pat. No. 5,048,524 to Bailey; andU.S. Pat. No. 5,066,859 to Karkar, additional detectors may be used tocompensate various measurement artifacts, particularly for variations inintensity of light entering the blood stream.

However, the correlation between blood parameters and data obtained byphotometric methods needs to be enhanced.

In U.S. Pat. No. 5,331,958 to Oppenheimer, a through photodetector and aremote photodetector are used for correction of light scattering due tosodium concentration fluctuations in the blood. Signals of the throughand the remote photodetector are amplified separately. The data from theremote detector is used to compensate or adjust the data of the throughdetector. These and other dual correction or compensation systems aresubject to the inherent unreliability of multiple detectors, as well ascalibration issues.

A disadvantage of other methods of photometric analysis of bloodproperties, which utilize conventional external tubing, includes adependence on the physical characteristics of the tubing such asdiameter, wall thickness, and light absorption factor.

Therefore, the need exists for a method and apparatus which require onlya single light detector, wherein the light scattering effect iseliminated or minimized. In such an apparatus, the mechanicalconstruction and data collection process is substantially streamlined ascompared to multiple photodetector devices. There is a need for a methodand apparatus having automatic calibration of the optical probesensitivity for each particular tubing that the probe is employed with;and further having a dilution technique applicable for determininghemodynamic parameters of a patient.

SUMMARY OF THE INVENTION

In accordance with the present invention, there is a novel method andapparatus for the photometric analysis of blood properties. The presentdesign allows for registration of blood optical density, provides signaladjustment for fluctuations in light scattering, which are related toblood electrolyte composition, as well as blood flow rate and otherblood parameters.

The present apparatus employs a single light emitter projecting a lightbeam along an illumination axis and a single receiving photodetector formeasuring light along a detection axis, wherein the illumination axis isnon colinear, offset or angled with respect to the detection axis.According to the method, the present invention includes projecting alight beam into a blood column along an illumination axis, and detectingthe light emerging from the blood column by a single receivingphotodetector receiving system along a detection axis that is noncolinear to the illumination axis.

The effects of light scattering are taken into consideration and havebeen eliminated by the location and orientation of the single receivingphotodetector detection axis with respect to the illumination axis. Itis understood the present system may be employed with additional sensorsor detectors, however these sensors are not required, or related to thelight scattering compensation, but are employed for different purposes.

Preferably a light emitter of isobestic wavelength is used to avoid theeffect of hemoglobin oxygen saturation.

The device may be used with an extracorporeal tubing system throughwhich blood flows, as well as being applied to patient' vessels and bodytissues which are capable of being transilluminated. Further, the devicemay be used during a dialysis session as well as surgeries, intensivecare procedures, measurements hemodynamic parameters of blood bydilution technique, or other blood property altering events.

The present method includes utilizing biologically natural indicators,as for example isotonic saline, glucose, or another solution, formeasurements by dilution technique and calibration procedures. There isno need in special markers, such as dye-green or the like, which stay ina patient's body for a long time.

In the drawings and in the detailed description of the invention thereare shown and described only principal embodiments of this invention andare of illustrative nature only, but not restrictive. Other embodimentsand technical realizations are applicable, all without departing fromthe scope and spirit of the invention.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a plot showing the light intensity changes caused by equaldiluting volumes of isotonic saline, water and hypertonic saline.

FIG. 2 is a plot showing the light intensity changes caused by the bloodflow rate changes.

FIG. 3 is a diagram showing the light scattering dependence onhematocrit level, blood electrolytes and blood flow rate.

FIG. 4a is a diagram showing a particular angle α in an emitted lightbeam that is characterized by stable light portion.

FIG. 4b is a diagram showing a principal dependence of absolute relativeerror of blood optical density measurement on an angle betweenIllumination Vector (IV) and Detection Vector (DV), where S_(NORM) issignal change while electrolyte composition remains unchanged, S_(HYPER)is signal change when electrolyte content was increased. An angle a isthe angle of zero error, and still some vicinity Δα is a range of angleswhere accuracy of measurement remains acceptable.

FIG. 5 is a plot showing registered light intensity signals for threebolus injections of equal volume of normal saline, water and 5% salinewith a sensor constructed and located to adjust for the scatteringeffect.

FIG. 6 is a plot showing the light intensity changes caused by the bloodflow rate changes registered with a sensor constructed and located toadjust for the scattering effect (lower trace) versus changes registeredwith traditional straight transilluminating sensor (upper trace).

FIG. 7 is a schematic of the present single emitter-single detectorsystem.

FIG. 8 is a schematic of an alternative embodiment of the present singleemitter-single detector system.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

Generally, the present method and apparatus include a probe 10 having ahousing 20, a light source 30 and a receiving photodetector 40 forcooperating with a length of light transmissive tube 50.

The light transmissive tube 50 may be any of a variety of conduitsincluding but not limited to a dialysis tubing system, an artery-veinextracorporeal tubing shunt, or blood vessel.

The housing 20 is sized to operably retain the light source 30 and thephotodetector 40. The housing 20 includes a light port 22 through lightfrom the light source 30 may pass. The housing 20 includes a tubereceiving passage 24 for receiving a length of the light transmissivetube 50, wherein which blood may flow through the tube. The housing 20is constructed to dispose the tube 50 intermediate the light port 22 andthe photodetector 40.

Referring to FIG. 7, the probe 10 embodying the present apparatus isshown in an illustrative form. The probe 10 includes the light source 30for generating a light beam. The generated light beam travels along anillumination vector IV. The illumination vector IV of the light may bedictated by the light source 30 itself, or by the configuration of thehousing 20. Specifically, the light port 22 of the housing 20 may beconfigured as a slit or aperture through which light from the lightsource 30 passes, and the housing thus permits only light passing alonga predetermined illumination vector IV to intersect the tubing 50. Asshown in FIG. 7, in a first embodiment the light source 30 and housing20 produce an illumination vector that is perpendicular to alongitudinal axis of the tubing 50.

To avoid the need in correction for oxygen saturation, a light source 30with an isobestic wavelength of emission spectrum may be used. The lightsource 30 is driven by light source driver 32. The light source driver32 may include optical feedback for source of supplied current, or maybe a dedicated laser diode driver.

The receiving photodetector 40 is spaced from the light port 22 tolocate a length of the light transmissive tube 50 therebetween. Thereceiving photodetector 40 is constructed to detect an intersection oflight along a detection vector DV. As blood flows through the tube 50, ablood column is exposed to the illumination vector IV and thephotodetector 40. Light initially entering the blood column along theillumination vector IV is absorbed and scattered in the blood flowingthrough the tubing 50. The receiving photodetector 40 is oriented toreceive only that part of the light passing through and emerging fromthe blood column that is scattered by the angle α with respect to theillumination vector IV. Thus, only light that is inclined from theillumination vector IV by the angle α reaches the receivingphotodetector 40. The detection vector DV of the receiving photodetector40 is offset from the illumination vector IV by the angle α, therebyallowing only light scattered by angle α to reach the receivingphotodetector. That is, the detection vector DV is inclined with respectto the illumination vector IV by the angle α. Although the receivingphotodetector 40 is shown as coincident with the detector port, it isunderstood the detector port may define the interface with the bloodcolumn wherein the receiving photodetector is spaced from the bloodcolumn. That is, the application of secondary optics may allow the lightsource 30 to be spaced from the light port 22 and the receivingphotodetector 40 to be spaced from the detector port.

Therefore, in the present design light is projected along a linearillumination vector IV and the detection vector DV of the receivingphotodetector 40 is offset from the illumination vector by the angle αsuch that only light scattered by the angle α intersects the receivingphotodetector.

The signal detected by the receiving photodetector 40 is amplified by anamplifier 42 to produce an amplified signal. The amplifier 42 may applya logarithmic transformation function to be able to further operate interms of optical density of the blood rather than merely detected lightintensity. Conveniently, the amplified signal is applied to ananalog-to-digital converter 44, and passes into a microprocessorinterface 46. After processing the amplified and digitized signal toproduce data, the microprocessor interface 46 causes the data to appearat a conventional monitor 48, such as a computer.

An alternative embodiment of the optical probe 10 is shown in FIG. 8. Inthis embodiment, the relative orientation between the light source 30and the receiving detector 40 is modified. However, the illuminationvector IV remains offset from the detection vector DV by the angle β. Itis understood that the angle between the illumination vector IV and thedetection vector DV, shown as α in FIG. 7 and β in FIG. 8 may not beidentical. While the angles β and β could be equal, they may varydepending upon the specific configuration of the system.

As shown in FIG. 8, the probe 10 is constructed so that the housing 20locates the light source 30 straight across from a receiving detector40. That is, the light port 22 and the photodetector 40 are located atthe same longitudinal position along the tubing 50, and areperpendicular to a longitudinal axis of the length of tubing as thetubing passes through the housing 20. In this configuration, theillumination vector IV is non perpendicular to the longitudinal axis ofthe tubing 50. That is, the illumination vector IV includes a componentthat extends along the longitudinal axis of the local length of thetubing 50 and hence blood column. In this configuration, the anglebetween the illumination vector IV and the detection vector DV is β.Locating the photodetector 40 and detection vector DV at the angle βfrom the illumination vector IV results in the photodetector registeringthe same portion of scattered light as the first configuration.

It is understood that some accommodation may be made for the defractionof the light as the light passes from the air through a wall of thetubing 50, into the blood, from the blood into a second wall of thetubing and from the tubing into the air. However, it is believed suchaccommodation is relatively small and may be accounted for by a smallshift in alignment of housing to correct diffraction by parallel layersof materials.

Theory

As stated, a change in hemoglobin concentration induces a change in theoptical density of blood. An amount of light passing through a bloodsample and emerging from the blood sample can be disposed andinterpreted to determine blood properties. However, there are severalfactors that influence light scattering in blood. These factors includeelectrolyte composition, flow rate and hematocrit level of the blood.These influences introduce error in hemoglobin evaluation as well asinjection bolus registration measurements made by a photodetector andlight source having aligned illumination vectors and detection vector(that is, the detector is displaced straightway across the tubing fromthe light source at a common longitudinal position and illuminationvector is collinear with the detection vector). Thus, in the systems ofthe prior art, the illumination vector and the detection vector arecolinear. These influences are conveniently referred to as the “scattereffect” and are illustrated in FIGS. 1, 2 and 3.

Referring to FIG. 4a, as it has now been discovered, there is aparticular angle α between the illumination vector IV of an emitted beamand a detection vector DV which is characterized by stable lightportion, scattered in this direction. For fixed distance between a lightsource (distance from the light port 22 to the receiving photodetector40), fixed blood column width and stable emitted light intensity, thesignal received by the receiving photodetector at the angle α from theillumination vector IV is a function of hemoglobin concentration ofblood and does not depend on light scattering.

Based on this understanding, the present optical probe 10 was designed.For this probe 10, the scatter effect is essentially eliminated due toselection of adequate angle as illustrated in FIG. 4b. Registereddetected light intensity signals for three bolus injections of equalvolume of normal saline, water, and 5% saline are shown in FIG. 5. Bloodflow related signals, shown in FIG. 6, also represents substantialimprovement.

Signal Processing

Signal processing for the evaluation of the blood optical properties,changes in liquid balance of the blood and cardiac output calculationare based on following:

Assuming the number of erythrocytes remains constant during blooddialyzing, changes in liquid content induce changes in blood hemoglobinconcentration. Relative volume change can be expressed as:$\begin{matrix}{\frac{\Delta \quad V}{V_{0}} = \left( {1 - \frac{C_{0}}{C}} \right)} & \left( {{Equation}\quad 1} \right)\end{matrix}$

where C₀ is the initial, or basic hemoglobin concentration; C is thecurrent hemoglobin concentration; V₀ is the initial volume and ΔV is thechange in volume. Concentration can be evaluated from knowing the amountof light emerged from the blood by applying Beer-Lambert Law:

I=I _(s) e ^(−κch)  (Equation 2)

where I is intensity of light being absorbed emerging from the bloodcolumn of a thickness h; I_(S) is light intensity of the light source; kis the absorption coefficient; and C is concentration of the medium.Relative volume change can be expressed through the light intensity as:$\begin{matrix}{\frac{\Delta \quad V}{V_{0}} = \left( {1 - \frac{\ln \quad \frac{I_{s}}{I_{0}}}{\ln \quad \frac{I_{s}}{I}}} \right)} & \left( {{Equation}\quad 3} \right)\end{matrix}$

where I_(S) is light intensity of the light beam; I₀ is the intensity ofemerged light for initial concentration of blood hemoglobin and I is theintensity of emerged light for current meaning of concentration.

Relative hemoglobin concentration change can be expressed through thelight intensity as: $\begin{matrix}{\frac{\Delta \quad C}{C_{0}} = \left( {1 - \frac{\ln \quad \frac{I}{I_{0}}}{\ln \quad \frac{I_{s}}{I_{0}}}} \right)} & \left( {{Equation}\quad 4} \right)\end{matrix}$

An important parameter of the system is the sensitivity of the probe 10.Different brands of tubing 50, even different sections of tubing of thesame brand may have different diameters and wall thickness. Anabsorption factor of the tubing material can also vary. Thus,sensitivity of the probe 10 should be known or determined for everyparticular tubing sample. It is also important to know any possiblechange in sensitivity of the probe 10 over relatively long periods ofmonitoring the blood properties.

A registered change in the optical signal received by the photodetector40 is proportional to a relative concentration change,$\frac{\Delta \quad C}{C_{0}},$

with a proportion coefficient, (a factor for sensitivity of the opticalprobe 10 to relative blood hemoglobin concentration change), K:$\begin{matrix}{\frac{\Delta \quad C}{C_{0}} = {{K \cdot \Delta}\quad U}} & \left( {{Equation}\quad 5} \right)\end{matrix}$

where ΔU is the change in the optical signal received by thephotodetector 40, measured in volts, for example. To determine K, aknown change is made in the relative hemoglobin concentrations and acorresponding signal change ΔU is registered.

Calibration

To provide correct data registering during long term monitoring of bloodproperties, measurements of hemodynamic parameters by dilutiontechnique, and determining different absolute blood parameters, anoptical sensor should be calibrated for particular operating conditions,such a current tubing system, initial blood parameters, as well asenvironmental changes occurring over the duration of the measurements.

When conventional hemodialysis is conducted, there are two major ways tocalibrate the system. The calibration includes producing a known changein a blood optical property and registering the corresponding change inthe measured (detected) signal.

In a first, more preferable method, the probe 10 is placed on the venousside of the dialysis tubing system, downstream from the dialyzer. In thefirst method of calibration, the ultrafiltration rate is changed by aknown amount, or completely terminated. When the ultrafiltration rate isaltered, the venous flow rate becomes greater by ultrafiltration ratechange Q_(UF). The change in the ultrafiltration rate produces adecreasing hemoglobin concentration.

Relative concentration changes in a venous line can be expressed throughblood flow rates as follows: $\begin{matrix}{\left. \frac{\Delta \quad C}{C_{0}} \right|_{V} = {- \frac{Q_{UF}}{Q_{B}}}} & \left( {{Equation}\quad 6} \right)\end{matrix}$

where Q_(B) is the known blood flow rate in an arterial side and Q_(UF)is the known ultrafiltration rate change. Thus, desired coefficient Kis: $\begin{matrix}{K = {{- \frac{Q_{UF}}{Q_{B}}}*\Delta \quad U_{UF}}} & \left( {{Equation}\quad 7} \right)\end{matrix}$

where ΔU_(UF) is change in the detected registered signal as caused bythe change in the ultrafiltration rate.

The relative change in hemoglobin concentration of the arterial blood iscalculated through a measured venous side concentration change accordingto: $\begin{matrix}{\left. \frac{\Delta \quad C}{C_{0}} \right|_{AR} = \left. {\left( {1 - \frac{Q_{UF}}{Q_{B}}} \right)\quad \frac{\Delta C}{C_{0}}} \right|_{V}} & \left( {{Equation}\quad 9} \right)\end{matrix}$

where $\left. \frac{\Delta \quad C}{C_{0}} \right|_{AR}$

is the relative change in arterial hemoglobin concentration,$\left. \frac{\Delta \quad C}{C_{0}} \right|_{V}$

is the relative change in venous hemoglobin concentration, Q_(UF) isultrafiltration rate, and Q_(B) is blood flow rate in the arterial side.

Dialysis machines of some brands turn ultrafiltration off duringdialysis session several times to conduct internal self-calibration.Advantageously, these self-calibrating procedures can be used forcalibration of the sensitivity optical probe to the relativeconcentration change.

Another method of calibrating the sensitivity of the probe is to employan indicator which dilutes the photometric density of the blood passingthrough the tubing 50. For example, an injection of an indicator, suchas normal saline or another conventional solution, is made into aninjection port of the dialysis system. The indicator injection should bemade upstream from the probe 10, so that all the indicator passesthrough the probe between the light port 22 and the photodetector 40.The indicator injection causes the optical property of the blood tochange, thus changing the light intersecting the photodetector 40 alongthe detection vector. The change in the amount of detected scatteredlight is registered by the probe 10, and specifically the photodetector40. Coefficient K can then be calculated through the volumes:$\begin{matrix}{K = {{- \frac{V_{Inj}}{V + V_{Inj}}}*\Delta \quad U_{Inj}}} & \left( {{Equation}\quad 8} \right)\end{matrix}$

where V_(Inj) is injection's volume; V=(Q_(B)−Q_(UF))*ΔT_(Inj), ΔT_(Inj)is the transit time of the injection bolus; and ΔU_(Inj) is integratedover ΔT_(Inj) registered signal change.

The calibration can also be made with the routine injection of adilution indicator into an injection port on an arterial side of thedialysis system. The probe 10 is then placed on arterial side,downstream from the place of injection. Again, all the injectedindicator must pass through the probe 10. No recalculation is requiredin this case.

A calculation of cardiac output (CO) of a patient can be made bydilution techniques with the following method.

By injecting a known amount of an indicator into a blood flow, thediluting effect of the indicator over a period of time can be accuratelydetermined by a sensor responsive to changes of photometric propertiesof the blood. The sensor is positioned so that the indicator passes thesensor, with the measured diluting effect being used to determinevarious blood parameters. The measurement may be made in anextracorporeal blood system in which optical measuring probes aresecured, for example, to tubing leading to exterior blood treatmentequipment such as hemodialysis machine, extracorporeal artery-veinpassive shunt, or the like. Thereafter, a bolus, or known volume, of anindicator material is injected into the bloodstream, and measurementsare made of changes in the photometric properties of the blood todetermine the passage of the bolus past the optical probe 10. Thechanges in such photometric properties of the blood can then be plottedand used to determine various hemodynamic parameters, particularlycardiac output of a patient.

With the aim of calibration for optical properties of different tubing,the above mentioned calibration injection of known amount of anindicator may be employed. The calibration injection must be madedirectly prior to the location of the probe, while simultaneouslymeasuring the blood flow through the blood system. A blood flowmeasuring device can be, for example, ultrasonic flow meter such asthose manufactured by Transonic Systems, Inc., Ithaca, N.Y. The cardiacoutput will be calculated as follows: $\begin{matrix}{{CO} = {\frac{V_{ind} \cdot S_{{cal} - {ind}}}{V_{{cal} - {ind}} \cdot S_{b - {ind}}} \cdot Q_{A - {cal}}}} & \left( {{Equation}\quad 13} \right)\end{matrix}$

where V_(cal-ind) is the volume of the calibrating injection,S_(cal-ind) is the dilution area under a curve generated by themeasurement of dilution produced by a calibrating injection.

Generally, cardiac output of a patient's heart may by calculated as:$\begin{matrix}{{CO} = \frac{K_{s} \cdot V_{ind}}{S_{b - {ind}}}} & \left( {{Equation}\quad 14} \right)\end{matrix}$

where V_(ind) is total volume of injected indicator; S_(b-ind) is thearea under a dilution curve representing the total optical densitychanges in the blood column over a time period; and K_(s) is thecalibration coefficient for a particular set of tubing and probe, asdetermined earlier.

As an advantage of using the probe for cardiac output measurement, itshould be noted the insensitivity of the probe to temperature variationsin blood, indicator, and the tubing carrying the blood. Further, sincethe probe can be secured in extracorporeal tubing system such as passivearterial-venous shunt or dialysis tubing, the present method of cardiacoutput measurement substantially reduces the invasive penetration into apatient's vital organs, such as the heart, as compared to Swan-Ganzcatheter.

It is understood the same calibration approach, including using a changein ultrafiltration rate or indicator injection can be used for sensorsother than optical sensors, wherein these alternative sensors canprovide a signal corresponding to any of a variety of measureable bloodparameters or properties including, but not limited to proteinconcentration, saline, electrolytes or any other measurable bloodproperty or parameters.

Therefore, any type of optical sensor, impedance, resistance orelectrical sensors which measure a changeable blood parameter such asthe sound or ultrasound velocity in blood can be calibrated. Electricalresistance of the blood can be measured, as the resistance depends onthe volume of red blood cells (hematocrit). Calibration can be providedfor ultrasound velocity sensors, as well as temperature sensors andoptical density, density or electrical impedance sensors can be used todetect changes in blood parameters.

A known change in the ultrafiltration rate will produce a correspondingmeasurable change in ultrasound velocity in the blood or if electricalresistance is measured, a voltage proportional to the ultrasoundvelocity. For example, having an initial venous blood flow Q_(b)=300ml/min, then after initiating ultrafiltration rate of Q_(f)=30 ml/min.is filtered out, the resulting change in concentration of largeparticles that were not withdrawn in the dialyzer (proteins, hematocrit)in the remaining flow 270 ml/min will be 100 (30 mil/min/300ml/min)=10%.

Thus, if in this example, the change in voltage of the ultrasound sensorwas 5 volts, the resulting correspondence provides that each 1 voltchange will represent 10%/5 volts or a 2% volume change. Thus, k=2%/1volt. As set forth above the blood volume can then be determined.

While a preferred embodiment of the invention has been shown anddescribed with particularity, it will be appreciated that variouschanges and modifications may suggest themselves to one having ordinaryskill in the art upon being apprised of the present invention. It isintended to encompass all such changes and modifications as fall withinthe scope and spirit of the appended claims.

What is claimed is:
 1. A method for measuring a blood volume change, themethod comprising: (a) connecting an arterial tubing portion of adialysis system to withdraw blood from a patient and connecting a venoustubing portion of the dialysis system to deliver filtered blood to thepatient; (b) operably coupling an optical sensor to the venous tubingportion of the dialysis system; (c) changing an ultrafiltration rate ofthe dialysis system; (d) determining at least one photometric propertyof blood at the venous tubing portion of dialysis system; (e)determining a calibration coefficient of at least the optical sensor;and (f) determining a blood volume change corresponding to thecalibration coefficient.
 2. The method of claim 1, wherein changing theultrafiltration rate includes changing the ultrafiltration rate by aknown amount.
 3. The method of claim 1, wherein changing theultrafiltration rate includes temporarily stopping the ultrafiltration.4. The method of claim 1, wherein determining the at least onephotometric property of blood includes determining a propertycorresponding to a protein in the blood.
 5. The method of claim 1,wherein determining the at least one photometric property of bloodincludes determining a property corresponding to a protein concentrationin the blood.
 6. The method of claim 1, wherein determining the at leastone photometric property of blood includes determining a propertycorresponding to a hematocrit of the blood.
 7. The method of claim 1,wherein determining the at least one photometric property of bloodincludes determining a property corresponding to red blood cells in theblood.
 8. The method of claim 1, wherein determining the at least onephotometric property of blood includes determining a propertycorresponding to an optical density of the blood.